In photoacoustic (PA) imaging, tissue irradiation using light (typically pulsed laser light), and subsequent thermo-elastic conversion of absorbed light to ultrasound, allows the detection of optically absorbing structures deep inside biological tissue with high resolution using ultrasound receive beamforming1 2. This technique is especially promising for functional imaging of the vasculature2, and of the blood oxygenation level using a multi optical wavelength approach based on the different optical absorption spectra of oxy- and deoxyhaemoglobin3. Photoacoustic imaging therefore holds promise for the diagnosis of vascular diseases and cancer4 and monitoring response to treatment5 6. In addition, gold nanoparticles, tailored to strongly absorb light in the NIR range, can serve as contrast media7 8, and their functionalisation for specific chemical targets allows early detection of e.g. cancer and atherosclerosis
Potentially, PA methods can provide an additional functional imaging modality, augmenting conventional ultrasound (US), for example in a real-time, safe, cheap, and versatile multimodal device for improved clinical diagnostics. For versatile imaging of the human body, an epi-style setup is preferred, combining the optical components with the acoustic probe for optical irradiation of the tissue from the same body surface as acoustic signal detection. In this way the influence of bones, acoustically attenuating tissue, and gas on ultrasound propagation from the illuminated tissue region to the acoustic probe can be reduced.
An important requirement for a clinically adequate and thus successful combination of PA and US imaging is an imaging depth of several centimetres, which is feasible in theory taking into account optical attenuation and transducer noise °. Such an imaging depth has, however, been difficult to achieve in practice. A reason is that the epiphotoacoustic setup causes severe clutter, which degrades contrast and limits imaging to depths considerably less than the noise-limited theoretical depth, at typically one centimetre or even less10 11. Clutter can emerge from strong PA transients that are generated at the site of tissue irradiation close to the ultrasound probe, where optically absorbing structures such as melanin and the microvasculature are exposed to the greatest intensity of irradiating light, or potentially elsewhere if a strong enough PA signal is generated. These transients obscure weak signals from deep inside the tissue when propagating directly to the acoustic receiver (direct clutter), and to the acoustic receiver via acoustic scattering from echogenic structures when propagating into the tissue (echo clutter).
Deep clinical PA imaging thus requires methods for clutter reduction to achieve the theoretical depth of several centimetres. For this purpose, displacement-compensated averaging (DCA) was previously developed11-13, a technique exploiting the clutter decorrelation that naturally occurs when palpating the tissue with the ultrasound probe, in motions parallel to the imaging plane. When compensating the resulting PA image sequence for the local relative tissue displacement, the “true” PA signal remains well registered and correlated whereas clutter decorrelates and can be reduced by averaging. DCA takes advantage of a combined PA and US system, because US speckle tracking provides the knowledge of local tissue displacement required for DCA.
Evaluation of DCA in combined PA and US imaging of human volunteers has demonstrated that clutter is an actual issue in clinical imaging, and that clutter reduction is feasible13. DCA, however, shows several disadvantages at the clinical application level. First, it can only be employed for easily palpable tissue such as breast and limb muscles, and requires a considerable amount of practice for controlled palpation in a free-hand approach. More significantly, however, is its limited clutter reduction and hence image contrast gain, determined by the maximum achievable tissue deformation on one side and by the minimum deformation required for clutter decorrelation on the other side. This typically results in a contrast gain not larger than three12, whereas a significantly larger contrast gain is desirable to achieve strongly increased imaging depth down to the noise limit.
Similar problems of clutter-limited image contrast exist in conventional US imaging, and potentially other forms of imaging such as optical coherence tomography (OCT). In conventional US echography, acoustic clutter may, for example, arise from acoustic scatterers interacting with side lobes or grating lobes, which may generate clutter echoes that return to the acoustic receiver either directly or after being scattered by other echogenic structures, or reverberation of ultrasound between acoustic scatterers that are proximal to the depth of interest13b. Approaches that are similar to DCA have been developed for US pulse-echo imaging independently and apparently without awareness of those developed for PA imaging, and perform with similar limitations to DCA in PA imaging13c. In OCT the strong and multiple optical scattering by tissue may generate substantial optical clutter. In the most common form of OCT this is substantially reduced by the use of a highly collimated beam of light13d. However, not only does this not fully remove the possibility of optical clutter generation at depths where the beam has been diffused by scattering, it has the substantial disadvantage that to produce an OCT image this beam must be scanned, reducing image frame rate. Alternative parallel acquisition methods using large area detectors offer potential for high frame and volume rate imaging but suffer from poor, optical clutter-limited, image contrast DCA methods may well be applicable to OCT and other imaging methods, although do not appear to have been tried and can be expected to perform with similar limitations to their performance in PA and US imaging.
Accordingly, for PA, US, OCT and other forms of imaging, an alternative approach to reducing or eliminating clutter is required.